1. Field of Invention
The present invention relates generally to the field of measuring magnetic fields. More specifically, the present invention is related to calibrating superconducting quantum interference device (SQUID) channels that are used in measuring magnetic fields.
2. Discussion of Prior Art
A SQUID magnetic sensor is at the heart of a sensitive magnetometer aimed at measuring magnetic fields below approximately 10xe2x88x9210 Tesla (T). This is the range of magnetic fields produced by living organisms (also called biomagnetic fields). For example, the human heart produces fields between 10xe2x88x9212 T and 10xe2x88x9210 T just outside of a chest surface. The magnetic fields emanated from the human brain, just outside of the head, are of the order of 10xe2x88x9214 T-10xe2x88x9212 T. These numbers can be compared with the earth""s magnetic field of about 10xe2x88x924 T and the typical urban magnetic noise of 10xe2x88x928 T-10xe2x88x926 T.
SQUIDs react to a magnetic flux rather than a field. Magnetic flux "PHgr"B is defined as the projection of the average magnetic field threading a given area along the area""s normal z, times that area A, or mathematically:
"PHgr"B=BZA
A low-Tc dc SQUID is an ultra-sensitive, low-noise transducer of magnetic flux "PHgr"B- to voltage, consisting of two nominally identical superconducting elements called Josephson junctions serially connected in a superconducting, electrically continuous loop. The SQUID loop is quite small in dimensions, typically 10xe2x88x924-10xe2x88x922 mm2. Today SQUIDs are typically produced on a chip, using Nbxe2x80x94Al junction technology, with junctions and the SQUID loop made of thin films. The micron-scale dimensions of the layout are defined using photolithographic techniques. The SQUID chip is typically enclosed in a superconducting shield screening the device from ambient magnetic flux. The magnetic flux to be measured is typically intercepted by considerably larger, 10-20 mm diameter loops or coils (called pick-up or detection coils) inductively coupled to a SQUID via an input coil. These coils are usually made of thin insulated superconducting (Niobium) wire wound over some non-conducting cylindrical support, although in some instances they are integrated on a chip with a SQUID. A single coil or a single loop intercepting magnetic field is called a magnetometer. More complex combinations of coils or loops, described in more detail below, form a gradiometer.
Since the SQUID and the coils must be kept in a superconducting state, they are immersed in liquid helium at temperatures only a few degrees above absolute zero (about xe2x88x92460xc2x0 F., or xe2x88x92269xc2x0 C., or 4xc2x0 K.). The double-wall vessel (space between walls being evacuated) intended for keeping and thermally isolating liquid helium is called a dewar. Dewars, in biomagnetic applications, are made largely of fiberglass in order to minimize magnetic interference with SQUIDs. Indeed, even non-magnetic metals are sources of secondary magnetic fields resulting from induced eddy currents.
Magnetocardiography (MCG) systems usually employ an array of sensors, for example 7 to 40. Measuring channel usually refers to one member of such an array, which comprises a single SQUID sensor inductively coupled to an arrangement of detection coils (magnetometer or gradiometer). Both SQUID and detection coils are typically mounted on a fiberglass support rod or on a fiberglass narrow, hollow cylinder. Shielded SQUID with its gradiometer together is usually called a sensor. The electrical leads or interconnects connect the sensor to associated electronic units stationed outside of a dewar at room temperature. Part of a channel that is physically attached to a fiberglass rod or a cylinder is called a probe. Additionally, the probes are essentially modular, so that each probe can be removed and inserted back into the dewar as necessary. Alternatively, all SQUID channels may be connected together in a common (non-modular) structure.
A response of a SQUID channel to a given input can be defined as a ratio of the output voltage to a combination of magnetic fields B="PHgr"B/A found at the detector coils. The form of this combination depends on a gradiometer type. For example, FIG. 1 illustrates a 2nd order symmetrical axial gradiometer consisting of three flat, axial, nominally identical coils or loops wound together. The loops contain 1-2-1 turns in the simplest implementation.
Because of the way the coils are wound, the supercurrents induced in the central loops flow in the direction opposite to the supercurrents in the two outer loops, so that the two outer coils produce signals of opposite polarity to the inner two-turn coil. Thus, this gradiometer produces a signal proportional to:
S2=BZ(z0)xe2x88x922Bz(z0+l)+Bz(z0+2l)
where Bz(z) is the z-component of magnetic field at a coordinate z, z0 is the coordinate of a lower detection coil, and l is the distance between neighboring coils called gradiometer""s base line, or base. In order to optimize signal-to-noise ratio (SNR), the base is chosen to be approximately equal to half of the distance from the lower detection coil to the magnetic field source (e.g., the heart). In gradiometers designed for heart measurements l is typically chosen to be about 5 cm, because the distance between the lower coil, placed about 2-3 cm above patient""s chest, and the heart is approximately 10 cm in a typical adult.
Similarly, one can wind a 3rd order gradiometer, which would consist in the simplest implementation of 1-2-2-1 loops, and so on, for even higher orders (see for example A. I. Braginski, H. J. Krause, and J. Vrba, in Handbook of Thin Film Devices, edited by M. H. Francombe, v. 3: Superconducting Film Devices, Chapter 6, p. 149, Academic Press (2000), incorporated here as a reference).
A gradiometer of k""s order acts almost as a magnetometer for nearby sources, while it subtracts spatially-constant magnetic field Bz and spatial derivatives up to order (kxe2x88x921): dBz/dz, dB2 z/dz2, etc. for distant sources. For example, a 2nd order gradiometer subtracts B and dBz/dz for distant sources. Thus, in this case the output voltage V divided by S2 can be considered to be the absolute channel response, in units of Volts/Tesla.
The output voltage V is the result of all electromagnetic processes taking place in the numerous electrical components of a given channel, including induced currents in the detector coils, induction coupling between the input coil and the SQUID, voltage response of a SQUID, filtering and electronic amplification of a signal, etc. Thus, the total channel response mixes SQUID""s transfer function with properties of the detection coils as well as with characteristics and settings of the associated SQUID electronics.
In order to determine absolute channel response V/S, one would need to know absolute values of the magnetic field at the positions of gradiometer detection coils. While this can be done, either by an actual measurement or by a calculation for a known field source, for the purposes of the specification, it is sufficient to find channel voltage response V alone, without dividing it by the magnetic signal S, as long as the field source is the same every time V is measured, as will be explained in more detail below. In what follows this voltage V is called the channel response.
As is clear, despite all possible precautions, nominally identical but physically different channels will inevitably present a certain spread of parameters. A number of factors contribute to channel-to-channel differences in the output response. Among these are geometrical and electronic factors.
Geometric factors are easily controlled. In the modular configuration, each probe is inserted into a specially designed space (notch) inside a dewar, so that its position and all the distances with respect to dewar""s overall geometrical shape (shell) and to the other probes are fixed and reproduced as well as mechanically possible; with proper care (such as making sure that there is no frozen air at the bottom of the notch) the reproducibility of a geometrical probe configuration is sufficiently good. This is even more so in the monolithic, non-modular system where all probes are rigidly connected together.
Electronic factors tend to have considerably more variation from channel to channel, and also as a function of time and system""s use. SQUID sensors may have different inductive couplings to their respective detector coils (gradiometers), and SQUIDs themselves may exhibit variability in their characteristics. Furthermore, the SQUID output signal is electronically amplified thousands of times. The multi-stage amplification with somewhat varying amplification coefficients produces a spread of the outputs, despite equal inputs.
Thus, for proper operation of an MCG system, different SQUID channels have to be calibrated in terms of their responses, to make them equal for equal input magnetic fields. To do that, one traditionally provides an arrangement in which each channel is exposed to a significantly non-uniform, nominally identical magnetic field, wherein the field source is preferably similar in its general type and distance from the sensor to the real source of interest (such as heart). Then, one measures channel voltage responses to this field. Inevitable differences in response will be found, for reasons outlined above. The system operator can then determine empirical ratios of channel responses, thus defining corresponding empirical coefficients for each channel with respect to one channel chosen as a reference. For example, the least sensitive (smallest response) channel number n is taken as such a reference, with its corresponding coefficient Cn taken to be unity, Cn=1. Any other channel, for example channel m, produces voltage response Vm such that Vm greater than Vn, with corresponding coefficient Cm=Vm/Vn greater than 1, where m is any channel index other then the index of the least sensitive channel n. In this case, division of the channel response voltages by corresponding coefficients, Vm/Cm=Vn makes the channel responses equal to the reference response of the least sensitive channel and to each other for equal fields at the detector coils. The result of this calibration can be also characterized as achieving equal sensitivity in system""s channels.
Once determined, these coefficients are introduced into the data-acquisition software; alternatively, amplification coefficients of each channel""s electronics may be adjusted by the factors Cm to achieve the same result.
As will be evident to the one skilled in the art, one can alternatively choose any channelxe2x80x94not necessarily the least sensitive onexe2x80x94as a reference channel. In general then some coefficients C will be greater than unity, while others will be smaller than unity. The chosen (reference) channel has coefficient equal to unity.
The method for finding these numerical coefficients (for initial system calibration purposes) in prior art systems consists of placing a standard circular current loop (or, alternatively, a matrix of such identical loops, which would be fed with identical currents one by one) under the dewar""s bottom, directly under each gradiometer""s lower detection coil. Additional care is taken to precisely position the dewar, with probes inside, over this field-producing loop or a matrix of loops, as shown in FIG. 2. FIG. 2 illustrates a dewar with an outer wall and an inner wall which enclose the probes 201 containing SQUID sensors (not shown) and gradiometer coils 202.
The current loop 203 is about the size of a gradiometer coil, i.e., about 20 mm in diameter. The standard current is supplied to the loop 203 from a calibrated current source 204, and the loop 203 is placed as precisely as possible under different probes 201, at a distance of about 10 cm from the lower detection coil, to mimic the heart""s field.
It should be noted that the vector of the dipole moment of such an xy-plane loop lies in z-direction, being perpendicular to the dewar""s bottom, while the dipole vector of a real heart lies typically in the xy-plane; in other words, the real heart looks more like a vertical rather than a horizontal loop. The loop 203 may be placed in the vertical (xz or yz) plane to better mimic the heart""s field configuration. This does not make too much of a difference for the purposes of the calibration procedure.
Next, the channel responses to nominally identical magnetic inputs induced by the loop 203 fed with standard current, said loop 203 placed in turn under each channel, are measured. Then, coefficients are calculated as explained above and, using these empirical coefficients, channel responses are equalized either through software or through hardware (electronics) adjustment.
This calibration method is somewhat cumbersome procedure, and it has its own sources of error, such as a problem of precise positioning of the current loop exactly under the respective gradiometers. This procedure is best done by the system manufacturer, in a shop equipped with means for precisely placing the loop and/or precise dewar positioning over it. Additionally, it is best performed in the absence of strong magnetic and RF interference, which can be achieved by having some magnetic shielding around the system, or by placing a system in a low magnetic interference environment. Such conditions may be hard to realize at the customer""s location, and they are generally best achieved in a specialized manufacturer""s shop.
Thus, one would ideally prefer to perform said calibration only once, in controlled environment. However, the problem is that a calibration will change in the course of time, as a result of any system modification or adjustment, and in particular as a result of any changes in electronics. Therefore, the calibration procedure is preferably repeated in prior art systems every time the system is moved, repaired, or an electronic part is replaced. These repeated, subsequent calibrations are typically performed by the customer under non-ideal conditions. This burdens the customer, and makes calibrations unreliable.
Whatever the precise merits, features and advantages of the above cited prior art systems, none of them achieve or fulfills the purposes of the present invention to provide an easy-to-use and accurate method and system of calibrating SQUID channels.
The present invention provides for a system and method for the improved calibration of magnetic field sensors based on superconducting quantum interference devices (SQUIDs). The improved calibration arrangement consists of a single fixed calibration ring encompassing all measuring channels, which is placed on an outer dewar wall, preferably near the middle of the gradiometer by height (i.e. approximately at the level of the middle set of gradiometer coils, or at a distance equal to one gradiometer base 1 above the level of the lower detection coils). During initial factory calibration, in addition to the small-loop procedure outlined above, either immediately subsequent to this procedure, or immediately preceding this procedure, a standard current is passed through the large calibration loop or ring, and channel responses U are measured and recorded. During re-calibration, the same standard current is passed through the ring, thus providing for the same magnetic field distribution over the gradiometer coils. Any channel responses that have changed from the original value U to a new value Uxe2x80x2, are identified and the corresponding empirical coefficients for these channels are changed from an old value of C to a new value Cxe2x80x2 given by:
Cxe2x80x2=(Uxe2x80x2/U)C
These new coefficients Cxe2x80x2 are then used to equalize channel responses as described above. To put it differently, they provide for equal sensitivity of different channels, which is the goal of the calibration procedure.
The system and method of the present invention are used to calibrate medical equipment utilizing SQUIDs used in the measure of magnetic fields associated with the heart.